Ultrasonic contrast agent detection and imaging by low frequency manipulation of high frequency scattering properties

ABSTRACT

A method for improved detection and imaging of ultrasound contrast agents using dual-band transmitted pulses, is described. The method is based on transmitting a pulse consisting of two frequency bands, a low frequency band which purpose is to manipulate the high frequency scattering properties of the contrast agent, and a high frequency band from which the image reconstruction is based. In addition, a general form of pulse subtraction is used to significantly suppress the received tissue signal.

RELATED APPLICATIONS

[0001] This application claims priority from U.S. Provisional PatentApplication Ser. No. 60/475,222 which was filed on May 30, 2003.

BACKGROUND OF THE INVENTION

[0002] 1. Field of the Invention

[0003] This invention relates to methods and systems for ultrasonicdetection and imaging of contrast agents located in soft tissue ortissue fluids.

[0004] 2. Description of the Related Art

[0005] Ultrasound contrast agents are typically made as solutions ofmicro gas bubbles or nano lipid particles. The gas bubbles typicallyshow strong and nonlinear scattering of the ultrasound, a phenomenonthat is used to differentiate the contrast agent signal from the tissuesignal. In the earliest applications (˜1985) the increased scatteringfrom the contrast agent within the transmitted frequency band was usedto enhance the scattering from blood. Later, second harmonic componentsin the nonlinearly scattered signal were used to further enhance thecontrast agent signal above the tissue signal in methods generallyreferred to as nonlinear contrast harmonic imaging.

[0006] The following two signal power ratios have vital importance forthe quality of performance of a contrast imaging system:

[0007] CTR—Contrast signal to Tissue signal Ratio. This gives the ratioof the signal power scattered from the contrast agent in a region to thesignal power scattered from the tissue in that region. This ratio isoften referred to as specificity.

[0008] CNR—Contrast signal to Noise Ratio. This gives the ratio of thesignal power scattered from the contrast agent in a region to the noisepower in that region. This ratio is often referred to as sensitivity.

[0009] The CNR determines the maximum depth for imaging the contrastagent while the CTR describes the enhancement of the contrast agentsignal above the tissue signal in the image and thus the capability ofdifferentiating contrast signal from tissue signal. High values of boththese ratios are therefore necessary for good imaging of the contrastagent.

[0010] The nonlinear distortion of the signal scattered from thecontrast agent is much stronger than for the tissue signal, a phenomenonthat is extensively used to enhance the CTR. In one method, receivedtissue signal components in the transmitted frequency band (linearcomponents) are reduced by combining the received signal from twotransmitted pulses with different amplitudes. In other methods, thesecond harmonic band of the nonlinearly scattered signal is obtainedeither by bandpass filtering or by combining the received signals fromtwo or more transmitted pulses with different polarities.

[0011] The contrast agent will typically undergo strong nonlinearoscillations with significant amount of energy scattered at higherharmonic components only if driven into oscillations well below itsresonance frequency and the harmonic component used for detection andimaging is often obtained in a bandpass filtering process. To obtaindistinct scattered harmonic components, the drive pulse typically has tobe relatively narrowbanded. The consequence of a relatively narrowbandedand low frequency drive pulse is the low image resolution typicallyobtained with harmonic imaging.

[0012] Also, the received nonlinear harmonic component from the contrastagent typically has low amplitude which reduces the CNR and may requireso high transmitted amplitude that the contrast agent bubbles aredestroyed. This can cause a problem when the inflow rate of contrastagent to the tissue region is low.

[0013] In addition, nonlinear contrast components scattered in theforward propagation direction will add in phase with the transmit fieldand hence accumulate. In tissue regions beyond a contrast filled area,these nonlinear contrast components may be linearly back-scattered fromthe tissue and falsely interpreted as contrast agent, hence reducing theCTR. Finally, a limitation in all methods based on nonlinear harmonicdetection is that nonlinear components in the tissue signal is preservedin the process, also limiting the CTR.

[0014] The new method described does not require nonlinear harmonicimaging and is therefore not constricted by the above mentionedlimitations encountered in nonlinear contrast harmonic imagingtechniques.

SUMMARY OF THE INVENTION

[0015] Ultrasound pulses containing both a low frequency band and a highfrequency band overlapping in the time domain, are transmitted towardsthe ultrasound contrast agent embedded in the tissue. The low frequencycomponents are used to manipulate the acoustic scattering properties ofthe contrast agent for frequency components in the transmitted highfrequency band, and the scattered bubble signal from the high frequencytransmitted components is used for image reconstruction. The lowfrequency components in the received signals can for example be removedthrough bandpass filtering of the signals around the high frequencyband.

[0016] The tissue signal is suppressed by transmitting at least two suchdual-band pulses for each radial image line with different phasesbetween the low and high frequency components, and performing a linearcombination of the back-scattered signals from the different pulses.

[0017] Due to nonlinear tissue elasticity, the transmitted low frequencypulse will slightly influence the wave propagation of the transmittedhigh frequency pulse resulting in slightly different high frequencysound speeds when altering the phase between the low and high frequencypulses. The resulting echoes may then have to be digitally interpolatedand adjusted relative to each other before combination to adequatelysuppress the high frequency tissue echoes.

[0018] With non-moving, temporary stationary tissue, one can for exampletransmit two pulses with different phase of the low frequency componentsand the same phase of the high frequency components, and perform alinear combination of the back-scattered signals from the two pulses.The scattered high frequency components from the contrast bubble will bemanipulated differently than from the tissue by the two low frequencypulses of different phases, and the bubble signal can be preserved whilethe tissue signal is heavily suppressed in the combination of the twoechoes.

[0019] When the tissue is moving, one may have to transmit more than twopulses for each radial image line to adequately suppress the receivedtissue signal. For example, one can transmit a set of M pulses, all withthe same phase of the high frequency components, but with differentphases of the low frequency components for each pulse. Theback-scattered signals from these pulses are combined in a pulse topulse high-pass filter as is commonly done in ultrasound imaging ofblood velocities to suppress the tissue signal. Typical filteringschemes that are used are FIR-type filters.

[0020] With electronic steering of the beam direction one would use thesame beam direction and focus for all the pulses that are combined tosuppress the tissue signal for each radial image line. Typical filteringschemes that are used are FIR-type filters or orthogonal decompositionusing for example Legendre polynomials, with filtering along the pulsenumber coordinate for each depth.

[0021] With mechanical scanning of the beam direction, as with annulararrays, one would typically transmit pulses with variations in the phaseof the low frequency components as the beam direction is sweptcontinuously, either applying a group of pulse number coordinates asinput to a FIR filter for each depth or with continuous inputs to a IIRfilter along the pulse number coordinate for each depth. The output ofthe FIR or IIR filter is then sampled for each depth and radial imageline to give the contrast agent signal, with attenuation of the tissuesignal, to be used for image reconstruction.

[0022] The present invention significantly increases the CNR relative toexisting methods by using the total scattered high frequency signal, andin particular the strong linear components, from the contrast agent andnot only nonlinear components of it.

[0023] Relative to nonlinear harmonic imaging methods, the presentinvention can use a more broadbanded transmit pulse and will henceachieve a higher range image resolution.

[0024] In addition, a higher transmit frequency can be used resulting ina significant increase in both lateral and range resolution relative tononlinear imaging methods.

[0025] Contrary to nonlinear imaging methods, the performance of thepresent invention will not be sensitive to the amplitude of the imagingpulses. Together with the indicated suppression of received tissuesignal with resulting increase in CNR, this facilitates nondestructivedetection and imaging of single contrast agent bubbles.

[0026] Other objects and features of the present invention will becomeapparent from the following detailed description considered inconjunction with the accompanying drawings. It is to be understood,however, that the drawings are designed solely for purposes ofillustration and not as a definition of the limits of the invention, forwhich reference should be made to the appended claims. It should befurther understood that the drawings are not necessarily drawn to scaleand that, unless otherwise indicated, they are merely intended toconceptually illustrate the structures and procedures described herein.

BRIEF DESCRIPTION OF THE DRAWINGS

[0027] In the drawings:

[0028]FIG. 1 displays the transfer functions from drive pressure toradial oscillation and to scattered pressure of a contrast bubbleundergoing small amplitude oscillations;

[0029]FIG. 2 illustrates transmit pulses containing both a low frequencypulse and a high frequency pulse where the high frequency pulse isplaced in the peak positive or peak negative period of the low frequencypulse;

[0030]FIG. 3 shows the radius responses from a bubble with resonancefrequency around 4 MHz when driven by the pressure pulses in FIG. 2;

[0031]FIG. 4 shows the far-field scattered pressure pulses from a bubblewith resonance frequency around 4 MHz when driven by the pressure pulsesin FIG. 2;

[0032]FIG. 5 depicts the absolute value of the Fourier Transform of thepressure pulses in FIG. 4;

[0033]FIG. 6 displays the result obtained by subtracting the twoscattered contrast pulses in FIG. 4 so that the high frequency tissuecomponents can be suppressed;

[0034]FIG. 7 illustrates the method of digital sampling rate increase(interpolation);

[0035]FIG. 8 shows a realization of the interpolation process by the useof polyphase filters; and

[0036]FIG. 9 shows schematically the adjustment and combination ofreceived echoes done in order to suppress the high frequency tissuecomponents.

DETAILED DESCRIPTION OF THE PRESENTLY PREFERRED EMBODIMENTS

[0037] The invention will now be described in more detail with referenceto the figures.

[0038] For small amplitude radius excursions, the mathematical equationsgoverning contrast bubble oscillation can be linearized and we obtainthe following transfer function from drive pressure to radial bubbledisplacement $\begin{matrix}{{H_{1}(\Omega)} = \frac{1}{\Omega^{2} - 1 - {i\quad \Omega \quad d}}} \\{{where}} \\{{d = \frac{b}{\omega_{0}m}},{\omega_{0}^{2} = \frac{s}{m}},{\Omega = \frac{\omega}{\omega_{0}}}}\end{matrix}$

[0039] Here, ω is the angular frequency and ω₀ is the resonancefrequency of the bubble while s is the stiffness of the gas and shell, mis the inertia of the surrounding liquid, and d is a damping factor ofthe resonant system.

[0040] The absolute value and phase angle of H₁(Ω) is shown in the upperand lower panel in FIG. 1a, respectively. In the lower panel, we seethat for drive frequencies well below resonance the displacement is πout of phase with the driving pressure. For frequencies well aboveresonance the bubble responds differently and the displacement and drivepressure are now in phase so that the bubble is increased in size whenthe drive pressure is positive and vice versa. Around resonance thedisplacement is approximately π/2 out of phase with the drive pressure.The absolute value of the amplitude of the transfer function is seen inthe upper panel of FIG. 1a. Going from frequencies below resonancetowards resonance the amplitude increases gradually culminating with aprominent peak around resonance for the situation with low damping(dashed line, d=0.1) and a considerable smaller peak for the situationwith higher damping (solid line, d=0.5). In both cases, the amplitudedecreases rapidly above resonance.

[0041] The transfer function from drive pressure to scattered pressurefor this small amplitude linearized situation is${H_{2}(\Omega)} = \frac{\Omega^{2}}{1 - \Omega^{2} + {i\quad \Omega \quad d}}$

[0042] The upper and lower panel of FIG. 1b display the absolute valueand phase angle of H₂(Ω), respectively. As previously, the dashed linesare results obtained setting the parameter d equal to 0.1 while thesolid lines are obtained for d equal to 0.5. The amplitude of thescattered pressure, as seen from the upper panel in FIG. 1b,significantly increases when going from drive frequencies belowresonance towards resonance. For drive frequencies above resonance, thescattered amplitude approaches a constant level. In the lower panel ofthe figure, we see that for drive frequencies well below resonance, thescattered pressure is in phase with the driving pressure. This meansthat the bubble oscillation is dominated by s, the stiffness of the gasand shell. For frequencies well above resonance, the bubble respondsdifferently and the oscillation is now dominated by m, the inertia ofthe co-oscillating fluid mass. The scattered pressure and drive pressureare now π out of phase. Around resonance the scattered pressure isapproximately π/2 out of phase with the drive pressure.

[0043] The purpose of the present invention is to heavily suppress thehigh frequency tissue echoes while maintaining the total high frequencycontrast agent echoes, and the essence of the invention is now describedby applying a simple two-pulse transmit scheme for each radial imageline.

[0044] In the first transmitted dual-band pulse, the high frequencycomponent 202 is placed in the positive peak of the low frequencycomponent 201 as shown in FIG. 2a, whereas in the second transmittedpulse, the high frequency component 213 is placed in the negative peakof the low frequency component 212 as shown in FIG. 2b. Notice that thetransmitted high frequency components (202 and 213) are identical andoccur at the exact same relative time from the pulse start. Thedifference between FIG. 2a and FIG. 2b is that the polarity of thetransmitted low frequency components are inverted with respect to eachother.

[0045]FIG. 3a shows the radius response from a contrast bubble withresonance frequency around 4 MHz when driven by the pulse in FIG. 2awhile the radius response from the same bubble when driven by the pulsein FIG. 2b is seen in FIG. 3b. The high frequency components (202 and213) in the transmitted pulses are here around 5 MHz and hence chosen tobe in the same area as the equilibrium resonance frequency of thecontrast bubble. This is, however, done only for purpose of illustrationand not a limitation in the present invention. From the bubble radiusoscillations, it is seen that the high frequency component in the firsttransmitted pulse occurs when the bubble is compressed (204) by the lowfrequency pulse, whereas the high frequency component in the secondtransmitted pulse occurs when the bubble is expanded (214) by the lowfrequency pulse. When compressed, the bubble will increase its resonancefrequency, while when expanded, it will reduce its resonance frequency.

[0046] From FIG. 1a we see that both the amplitude and phase angle ofthe radius oscillation will change for a given drive frequency whenchanging the resonance frequency of the bubble.

[0047] The resulting far-field scattered pressure from the contrastbubble when driven by the pressure pulse in FIG. 2a is depicted in timedomain in FIG. 4a and in frequency domain in FIG. 5a, while thescattered pressure obtained when driven by the pulse in FIG. 2b isdepicted in time domain in FIG. 4b and in frequency domain in FIG. 5b.The scattered high frequency fundamental component (209) in FIG. 5a issomewhat weaker than the scattered high frequency fundamental component(220) in FIG. 5b. Nonlinear scattered high frequency components (210 and221) are also somewhat different. Scattered low frequency components(207 and 218) have low amplitude and are not meant to be used for imagereconstruction. The purpose of the low frequency components is only tomanipulate the scattering properties of the contrast agent, i.e. to makethe bubble oscillate with such a low frequency that high frequencycomponents can be used to interrogate it while manipulated by the lowfrequency pulses.

[0048] From FIG. 1b we notice that both the amplitude and phase angle ofthe scattered pressure will vary for a given drive frequency whenvarying the resonance frequency as done when manipulating the bubble bythe low frequency pulses.

[0049] To suppress the tissue signal the two scattered pressure pulsesin FIG. 4 are then subtracted and the result is depicted in time domainin FIG. 6a. In FIG. 6b, we see the spectrum of the resulting pulse. Wenotice that even if the two high frequency drive pulses in FIG. 2 occurat the exact same relative time from the pulse start, due to themanipulation by the low frequency pulses, the scattered high frequencyenergy from the bubble is not canceled or significantly reduced in thesubtraction process of the two bubble echoes. We may thus utilize thetotal scattered high frequency energy from the bubble for imagereconstruction and not only a nonlinear component of it as done in allnon-destructive nonlinear contrast agent detection techniques.

[0050] It is also possible to use two transmit pulses where one of thetransmit pulses only contains the high frequency imaging pulse whereasthe other transmit pulse contains both the manipulating low frequencypulse and the high frequency imaging pulse overlapping in the timedomain. This would then be a version of amplitude modulation for the lowfrequency transmit pulse.

[0051] When the tissue is moving, it may be advantageously to transmitmore than two pulses for each radial image line to adequately suppressthe received high frequency tissue signal. For example, one can transmita set of M pulses, all with the same phase of the high frequencycomponents, but with different phases of the low frequency componentsfor each pulse. The back-scattered signals from these pulses arecombined in a pulse to pulse high-pass filter as is commonly done inultrasound imaging of blood velocities to suppress the tissue signal.

[0052] With electronic steering of the beam direction one would use thesame beam direction and focus for all the pulses that are combined tosuppress the tissue signal for each radial image line. Typical filteringschemes that are used are FIR-type filters or orthogonal decompositionusing for example Legendre polynomials, with filtering along the pulsenumber coordinate for each depth.

[0053] With mechanical scanning of the beam direction, as with annulararrays, one would typically transmit pulses with variations in the phaseof the low frequency components as the beam direction is sweptcontinuously, either applying a group of pulse number coordinates asinput to a FIR filter for each depth or with continuous inputs to a IIRfilter along the pulse number coordinate for each depth. The output ofthe FIR or IIR filter is then sampled for each depth and radial imageline to give the contrast agent signal, with attenuation of the tissuesignal, to be used for image reconstruction.

[0054] Acoustic wave propagation is in the linear regime governed by thelinear wave equation where the speed of sound is defined as$c_{0} = \frac{1}{\sqrt{\rho\kappa}}$

[0055] where ρ is the density and κ is the compressibility of thepropagation medium. Due to nonlinear tissue elasticity we get, based ona plane wave approximation, a nonlinear propagation velocity thatdepends on the wave pressure as

c=c ₀{square root}{square root over (1+2βκρ−2β²(κρ)²)}≈c ₀(1+βκρ)

[0056] where β is a nonlinearity parameter accounting for nonlinearintermolecular forces in the propagation medium and β is the acousticpressure. The last approximation is valid for the case when κρ<<1. Inmedical ultrasound imaging, κρ typically lies in the range from2·10^(−3 to) 2·10⁻⁵ whereas, β is around 5.

[0057] Ultrasound wave propagation in tissue is hence a weak nonlinearprocess for intensities commonly applied in medical imaging. Due to thenonlinear tissue elasticity, wave propagation of the two dual-bandpulses displayed in FIG. 2 will be slightly different. The highfrequency component (202) in FIG. 2a, occurring during the positivepressure swing of the low frequency component, will travel with aslightly higher sound speed than the high frequency component (213) inFIG. 2b, occurring during the negative pressure swing of the lowfrequency component. Using Eq.12 with c₀ equal to 1500 m/s, we get atypical sound speed of 1500.5 m/s for the transmitted high frequencycomponent in FIG. 2a and 1499.5 m/s for the high frequency component inFIG. 2b.

[0058] The consequence of this difference in sound speed is that the tworesulting high frequency echoes obtained from the indicated transmitpulses may have to be slightly time-shifted relative to each otherbefore combined to adequately suppress the tissue echoes. Thesetime-shifts are typically smaller than the sampling interval of thesignal which requires interpolation for adequately accurate signalvalues.

[0059]FIG. 7 displays schematically an interpolation method of samplingrate increase by a factor of I. The interpolation is here done by firstintroducing I-1 zeros between each sample in the original sequence x(n)with a sampling rate of F_(x) to obtain the desired sampling rateIF_(x).

[0060] Mathematically, the resulting sequence can be described as$\begin{matrix}{{v(m)} = {\sum\limits_{m = {- \infty}}^{\infty}\quad {x\left( {m/I} \right)}}} & {for} & {{m = 0},{\pm I},{{\pm 2}I},}\end{matrix}$

[0061] and v(m)=0 otherwise

[0062] The sequence v(m) is then passed through a lowpass filter h(m) toobtain the desired output y(m) and this lowpass filter is typicallyimplemented as a linear phase FIR-type filter where the z-transform ofthe filter is defined as${H(z)} = {\sum\limits_{k = 0}^{M - 1}\quad {{h(k)}z^{- k}}}$

[0063] for a filter of length M.

[0064] The direct-form realization of the interpolation algorithm iscomputationally not very efficient due to all the zeros in the sequencev(m). To increase the efficiency, the filter h(m) can be divided into aset of smaller filters of length K=M/I, where M is selected to be amultiple of I. This set of smaller filters are usually called polyphasefilters and have unit sample responses

g _(k)(m)=h(k+mI) for k=0,1, . . . , I−1 and m=0,1, . . . , K−1

[0065] Thus, the polyphase filters perform the computations at theoriginal low sampling rate F_(x) and the rate conversion results fromthe fact that I output samples are generated, one from each of thefilters, for each input sample. Interpolation by use of polyphasefilters are shown schematically in FIG. 8. Here, the polyphase filtersare arranged as a parallel realization and the output of each filter isselected by a commutator rotating in the counterclockwise direction. Thedecomposition of h(m) into the set of I subfilters with impulseresponses g_(k) (m) results in filtering of the input samples x(n) by aperiodically time-varying linear filter g(m,k).

[0066]FIG. 9 shows schematically the adjustment and combination processdone to suppress the received high frequency tissue signal components.The received echoes are first temporally interpolated and given avariable time adjustment (901) before the pulse to pulse combination(902) to heavily suppress the high frequency tissue components.

[0067] The separated contrast agent image is typically shown as anoverlay with different color or pattern of the standard tissue image asobtained with only one of the transmit pulses for each radial imageline.

[0068] If a weak tissue background is desired in the ultrasound image,the interpolation and time adjustment process may not be necessary, orthe time adjustment can be set to allow for a part of the obtained highfrequency tissue signal to remain after the combination. The resultingcontrast agent image will then be shown as an overlay on the resultingtissue image with a different intensity.

[0069] Thus, while there have shown and described and pointed outfundamental novel features of the invention as applied to a preferredembodiment thereof, it will be understood that various omissions andsubstitutions and changes in the form and details of the devicesillustrated, and in their operation, may be made by those skilled in theart without departing from the spirit of the invention. For example, itis expressly intended that all combinations of those elements and/ormethod steps which perform substantially the same function insubstantially the same way to achieve the same results are within thescope of the invention. Moreover, it should be recognized thatstructures and/or elements and/or method steps shown and/or described inconnection with any disclosed form or embodiment of the invention may beincorporated in any other disclosed or described or suggested form orembodiment as a general matter of design choice. It is the intention,therefore, to be limited only as indicated by the scope of the claimsappended hereto.

We claim:
 1. A method for detection and imaging of ultrasound contrastagent in a region of soft tissue, where, at least one ultrasound pulsecontaining frequency components in separated low and high frequencybands, is transmitted toward said region, the time signals of said lowand high frequency components are overlapping in time, so that saidtransmitted low frequency components are used to manipulate the acousticscattering properties of the contrast agent for the incident frequencycomponents in said high frequency band, where the scattered signal fromsaid transmitted high frequency components is used for ultrasound imagereconstruction of said region.
 2. A method for detection and imaging ofultrasound contrast agent according to claim 1, where to reduce the lowfrequency components in the received signal before image reconstruction,the received signal is filtered in range to attenuate the low frequencycomponents.
 3. A method for detection and imaging of ultrasound contrastagent according to claim 1, where, at least two ultrasound pulses aretransmitted consecutively toward said region with substantially samefocus and direction of the ultrasound beam, the phase of said lowfrequency components in said pulses is changed for each transmittedpulse, the received signal from the consecutive pulses are for eachdepth combined to suppress tissue signal components.
 4. A method fordetection and imaging of ultrasound contrast agent according to claim 3,where, the ultrasound beam direction is scanned in steps, and for eachradial image line a group of pulses is transmitted, each pulse withdifferences in the phase and/or amplitude of said low frequencycomponents, the back-scattered signal is sampled as a function of depthfor each pulse and the samples for each depth are combined linearlyalong the pulse number coordinate, the linear combination for each depthbeing selected so that the received high frequency components from thetissue are heavily suppressed, while the received high frequencycomponents from the contrast agent are not significantly suppressed bythe linear combination, so that the output of the linear combination isused to form the image of the contrast agent for each depth along saidradial image line.
 5. A method for imaging and detection of ultrasoundcontrast agent according to claim 1, where, multiple ultrasound pulsesare transmitted consecutively while the ultrasound beam direction isswept continuously over the region to be imaged, the phase and/oramplitude of said low frequency components of the transmitted pulsesbeing changed for each transmitted pulse, the received signal beingsampled in depth along the beam direction, and the samples used as inputto high-pass filters operating along the pulse number coordinate foreach depth, the output of the high-pass filters for each depth beingsampled at defined radial image lines to be used for construction of thecontrast agent image as a function of depth along said radial imagelines.
 6. A method for imaging and detection of ultrasound contrastagent according to claim 3, where the received signals from theconsecutive pulses are digitally interpolated and temporally adjustedrelative to each other before combination to adequately suppress thehigh frequency tissue components.
 7. A method for imaging and detectionof ultrasound contrast agent according to claim 3, where in saidtransmitted pulses the amplitude of said high frequency bands may beseparately adjusted.
 8. A method for imaging and detection of ultrasoundcontrast agent according to claim 3, where in said transmitted pulsesthe amplitude of said low frequency bands may be separately adjusted. 9.A method for imaging and detection of ultrasound contrast agentaccording to claim 1, where the resulting contrast agent image is shownas an overlay on the resulting tissue image with a different intensity,color, or pattern.
 10. A method for imaging and detection of ultrasoundcontrast agent according to claim 6, where the temporal adjustment isdone so that a part of the received high frequency tissue signal remainsafter the combination and is shown as a weak tissue image overlayed bythe contrast agent image.
 11. A method for imaging and detection ofultrasound contrast agent according to claim 5, where the receivedsignals from the consecutive pulses are digitally interpolated andtemporally adjusted relative to each other before combination toadequately suppress the high frequency tissue components.
 12. A methodfor imaging and detection of ultrasound contrast agent according toclaim 5, where in said transmitted pulses the amplitude of said highfrequency bands may be separately adjusted.
 13. A method for imaging anddetection of ultrasound contrast agent according to claim 5, where insaid transmitted pulses the amplitude of said low frequency bands may beseparately adjusted.